In vitro Mechanical Testing of Glass Fiber-reinforced
Composite Used as Dental Implants
Abstract
Aim: The aim of this study was to evaluate the design of fiber-reinforced composite (FRC) on some mechanical
properties of a dental implant.
Methods and Materials: FRC implants were fabricated using different polymerization conditions and designs
of the glass-fiber structure. Specimens were tested with a cantilever bending test and a torsional test. The
degree of monomer conversion (DC%) was measured using a Fourier transform infrared spectroscopy (FTIR).
Results: Statistical analysis showed significant differences between groups revealing mean fracture load
values from 437 N to 1461 N. The mean torsional force in fracture varied from 0.01 to 1.66 Nm. The DC%
varied from 50% to 90%.
Conclusion: This study suggests by modifying the polymerization conditions and fiber orientation of FRC
implants, the biomechanical properties of an FRC can be tailored to the needs of dental implants.
Keywords: Dental implant, fiber-reinforced composite, FRC, mechanical properties, bending overload, torsion
load, failure force
Citation: Ballo AM, Lassila LV, Närhi TO, Vallittu PK. In vitro
o Mechanical Testing of Glass Fiber-reinforced
Composite Used as Dental Implants. J Contemp Dent Pract 2008 February;(9)2:041-048.
© Seer Publishing
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The Journal of Contemporary Dental Practice, Volume 9, No. 2, February 1, 2008
Introduction
Dental implant treatment is one of the
commonly used conservative methods of
permanently replacing one or more missing
teeth.1 Dental implants have traditionally been
made of metal. The use of pure titanium or a
titanium alloy has been the material of choice.
Occasionally complications during the first year
of functional loading occur, but the most common
complications are loosening of abutments or
bridge screws and esthetics related.2 Functional
occlusal load and overloading of dental implants
during mastication has been extensively studied
and discussed,3,4 but less work has been done on
torsional loading forces of the implants.5 In an
oral environment dental implants are subjected
4,6,7
to multidirectional forces.
Forces can have
either vertical or horizontal components, but also
4
inclined and torsional forces may exist. The
bending moment due to forces deviating from the
direction of the implant axis can produce higher
stress in the implant and implant-bone interface
than a direct axial load.4,8
The magnitude of occlusal forces varies
depending on tooth position in the mouth.4 The
mean maximal occlusal forces in the incisal
area have been reported to range from 264 N to
370 N11,12 whereas the maximum biting force in
the posterior area can be around 800 N.13
Fiber reinforced composites (FRC) are a relatively
new group of materials investigated in dental
or medical applications over the last 30 years.14
Their use is growing in many dental applications
including use in implant-supported prostheses.15-17
In the FRC the fibers are made of E-glass or
quartz and are translucent which provides for
18
favorable esthetics. Microbial adhesion to
FRC has been shown to be lower in glass FRC
19
than in a FRC made of polyethylene fibers.
Furthermore, an FRC implant is relatively easy to
grind and modify directly in the mouth in order to
properly restore the implant abutment area with
composite superstructures without having a risk
of overheating underlying bone.
Proper treatment planning for implant placement
and careful crown fabrication with optimal
cusp inclination can normally prevent implant
overload.9 The excessive bending moments may
cause stress concentration and micro-fractures in
alveolar bone and even implant fractures.
Mechanical properties of FRCs are well
documented by many investigators.20,21 The
optimal mechanical properties for FRC in bending
can be obtained with continuous unidirectional
fibers.22 The modulus of elasticity of FRC (20-40
GPa) varies according to fiber volume and can
be tailored to simulate stiffness close to natural
bone.23
In addition, a mismatch of the modulus of
elasticity of the bone and the high stiffness
metallic implant induces the stress-shielding
effect, which means the implant does not strain
and load the bone physiologically similarly as
the absence of the implant. This causes loss of
10
bone and bone contact. For this reason, the
concept of using a material with similar modulus
of elasticity to bone has become an issue of
interest. By lowering the modulus of elasticity,
the mechanical properties of the implant can
better match the properties of bone which further
provides mechanical stimulus to the bone.
Effective wetting of fibers by the resin matrix,
also referred to as resin impregnation, is a
pre-requisite for their effective use.20,24 In the
case of photopolymerization of FRC, the light
intensity, exposure time, and the polymerization
temperature has an effect on flexural properties
25
and monomer conversion. Effects of water
sorption on the flexural properties of FRCs have
also been reported.26,27 Recently, attempts have
been made to use FRC as implant material in
head and neck surgery.28 Our previous study
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The Journal of Contemporary Dental Practice, Volume 9, No. 2, February 1, 2008
showed preosteoblasts can mature and proliferate
well on the FRC substrates giving support to the
use of FRC in a bone environment.29
prepare the specimens. The size of specimens
was based on the assumption if they were to
be used as oral implants their load-bearing
capacity should exceed average maximum
occlusal forces within the physiological strain
limit of bone.33 The fibers were inserted into the
mold along the long axis of the specimens. Both
continuous unidirectional resin preimpregnated
(prepregs) and manually impregnated resin
fibers were used (Table 2). The resin system
was a dimethacrylate-monomethacrylate
(BisGMA-PMMA) system that produced a semiinterpenetrating polymer network to form a
polymer matrix.
Bioactive glass (BAG) implant coatings have
been shown to improve osseointegration both in
in vitro
o and in vivo
o conditions.30,31 BAG particles
may be used as a bioactive component in FRC
devices since they can be embedded within the
resin matrix or applied on the surface of the FRC
implants which may improve osseointegration.
The aim of this in vitro
o study was to evaluate
bending and torsion properties and degree of
monomer conversion (DC%) of FRC implant
shaped test specimens in order to determine their
suitability as a potential implant material.
Experimental materials and polymerization
conditions of the specimens are summarized in
Table 2. Six groups of specimens were prepared
with eight specimens fabricated for each group.
Various polymerization conditions were tested for
optimizing the polymerization of the resin matrix.
The test specimens were polymerized with an
Optilux 501 hand light-curing unit (Kerr Mfg.,
Orange, CA, USA), in a light curing oven at 80°C
(LicuLite, Dentsply De Trey GmbH, Dreieich,
Germany) and post-cured at 120°C which was
the temperature of glass transition (Tg) of the
pBisGMA-pTEGDMA-copolymer.
Methods and Materials
The materials used for the fabrication of the
specimens for bending and torsion tests are listed
in Table 1. Each specimen contained seven fiber
reinforcements each of them consisting of 4000
single glass fibers of 15 μm in diameter. The
E-glass fiber composition was: 55% SiO2, 15%
Al2O3, 22% CaO, 6% B2O3 and 0.5% MgO, >1.0%
Fe +Na + K.
The modulus of elasticity of FRC material is
mostly affected by the volume of fiber and the
degree of conversion of the polymer matrix.32 In
this unidirectional FRC the modulus of elasticity is
between 20 and 40 GPa.
After polymerization, the specimens were wet
ground with 1200 grit (FEPA) silicon carbide
grinding paper after which the diameter and
length of the specimens were measured at
three different places in order to record the final
dimensions of the specimens. The specimens
were conditioned in air at room temperature for
two days before mechanical testing.
Specimen Preparation
Molds with an internal diameter of 4 mm and
length of 20 mm and 30 mm were used to
Table 1. Materials used in the study.
*Bis-GMA, bisphenol A-glycidyl dimethacrylate.
**TEGDMA, triethylenglycoldimethacrylate.
***PMMA, poly methyl methacrylate, Mw 220.000
**** E-glass, electrical glass
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Table 2. Classification of experimental FRC specimens used in this study
according to the polymerization conditions and design of fiber reinforcement.
Note: everStick™ was used in the fabrication of PMMA specimens in groups A, B, C, D, and E whereas in
group F the reinforcement was done using continuous unidirectional E-glass with manual bisGMA-TEGDMA
resin impregnation (Stick Resin, StickTech).
* see Figure 6.
Specimen Testing
For the cantilever bending test, 10 mm of each
specimen (4 mm diameter and 20 mm long) was
embedded from one end into the center of a 10
mm x 10 mm x 10 mm acrylic resin block. The
cantilever bending test was performed to measure
the flexural properties of the specimens. A jig
was used to provide a 45° angle between the long
axis of the implant and the direction of the loading
force (Figure 1).
The specimens were loaded to failure using a
Lloyd Model LRX material testing machine (Lloyd
Instruments Ltd., Fareham, England) with acrosshead speed of 1.0 mm/min until fracture. Fracture
force was determined as an audible crack or 10%
reduction in force indicated by the testing device.
The peak force of the failure of each specimen
was recorded with Nexygen PC software (Lloyd
Instruments Ltd., Fareham, England).
Figure 1. Cantilever bending test of the
specimen.
into the center of a 10 mm x 10 mm x 10 mm
acrylic resin block. Specimens were tested
for the resistance of torsional force to failure
using the same Lloyd Model LRX universal
testing machine but with an angle speed
of 16.2 degrees/min. A TP-2KMCB torsion
sensor (KYOWA Machine Ltd., Tokyo, Japan)
was used to measure maximum torsional
force of the specimen (Figure 2).
Failure types of test specimens after the bending
test were visually classified to the following
groups:
A. Delamination of the prepregs.
B. Fracture of the test specimen.
C. Delamination of the fiber weaves from the
core.
D. For the torsional force test, 10 mm of each
cylindrical specimen, 4 mm diameter and 30
mm long, were embedded from both ends
SEM micrographs illustrate the delamination of
prepregs and fiber weaves (Figures 3 and 4).
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The Journal of Contemporary Dental Practice, Volume 9, No. 2, February 1, 2008
The DC% of the polymer matrix of mechanically
tested specimens after various polymerization
modes was monitored using Fourier transform
infrared spectroscopy (FTIR) (Spectrum
One, Perkin Elmer, Beaconsfield Bucks, UK)
with an attenuated total reflectance (ATR)
sampling accessory. The mechanically tested
specimens were ground with a grinding stone
to collect powder to be placed on the surface
of the detector ZnSe-ATR crystal of the FTIR
spectroscope. The spectrum of unpolymerized
resin matrix of the everStick reinforcement (Stick
Tech Ltd, Turku, Finland) was used to measure
the DC% of polymer powder of each specimen.
Each spectrum was recorded with 16 scans using
a resolution of 4 cm. The DC% was calculated
from the aliphatic C=C peak at 1638 cm and
normalized against the aromatic C=C peak at
1608 cm according to following formula:
Figure 2. Schematic diagram of torsional loading test
of the specimen.
Where:
Caliphatic = absorption peak at 1638 cm−1
of the cured specimen
Caromatic = absorption peak at 1608 cm−1
of the cured specimen
Ualiphatic = absorption peak at 1638 cm−1
of the uncured specimen
Uaromatic = absorption peak at 1608 cm−1
of the uncured specimen
Figure 3. SEM micrographs illustrate the delamination
of the prepregs.
Data Analysis
Data were analyzed statistically with analysis of
variance (ANOVA) (SPSS, SPSS Inc., Chicago,
IL, USA) followed by a Tukey’s post hoc analysis
using a significance level of p<0.05.
Results
Tables 3 and 4 illustrate the mean fracture
load values and standard deviations for the
specimens. The ANOVA revealed post-curing
significantly affected the fracture load (p<0.001).
The mean fracture load values were: 437 N for
group A, 581 N for group B, 961 N for group C,
1108 N for group D, 1461 N for group E, and
1200 N for group F. The torque could not be
determined for the specimens in groups A, B, C,
and D because of delamination of FRC prepregs
in the early stage of loading. The mean torque
Figure 4. SEM micrographs illustrate fracture and
delamination of the fiber weaves.
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The Journal of Contemporary Dental Practice, Volume 9, No. 2, February 1, 2008
Table 3. The mean failure load values and standard deviation of the test
specimens in cantilever bending test (n=8). Same superscript letter indicates
groups did not differ statistically significantly (p<0.05).
Table 4. The mean maximal torsional force and standard deviation of the test specimens (n=8).
Same superscript letter indicates groups did not differ statistically significantly (p<0.05).
Table 5. The degree of conversion (DC%) and standard deviation of polymer matrix
of tested specimens after different polymerization conditions. Same superscript
letter indicates groups did not differ statistically significantly (p<0.05).
was 1.66 Nm (range 1.3 to 1.9 Nm) for group E
and 1.0 Nm (range 0.8 to 1.4 Nm) for group F.
Three types of failure patterns were found
(Figure 5). Failure pattern A was found in groups
A-D specimens, B in group E, and C in group F.
The DC% is presented in Table 5. The mean
DC% increased in all groups by increasing the
time of light curing and by increasing the postcuring temperature. The specimens in group A
had the lowest DC%; the mean DC% was 53%
for group A, 61% for group B, 73% for group C,
75% for group D, and 90% for group E and F.
Discussion
The mechanical link between a natural tooth
and the surrounding bone is provided by the
periodontal ligament (PDL) which controls
the dynamics of the tooth under load. By
comparison, a dental implant is integrated to the
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The Journal of Contemporary Dental Practice, Volume 9, No. 2, February 1, 2008
Figure 5. Schematic diagram of failure patterns of test specimens after
cantilever bending test in Group A-F:
A. Delamination of the prepregs.
B. Fracture of the test specimens.
C. Delamination of the fiber weaves from the core.
Table 6. Comparison of mechanical properties of different implant materials.
bone without a PDL. The mismatch of stiffness
between bone and implant material can lead to
implant failures especially in conditions where
either bone quantity or quality is poor. This can
occur if the tensile or compressive load exceeds
the physiological limit of bone tolerance and
causes microfracture at the bone-to-implant
interface or initiates bone resorption. On the other
hand, implant material that is too stiff does not
provide the normal physiological stimulus on bone
which causes loss of bone around the implant.
dental implant material. Recently, Zirconia has
been introduced as a new ceramic dental implant
material, but has a stiffness (modulus 200 GPa)
that is too high compared to bone. It also has
a very high modulus of elasticity (17–24 GPa)
compared to human bone.
As a metal substitute, Zirconia possesses
good chemical and physical properties, like
low corrosion potential and low thermal
conductivity.39-41 Furthermore, its biocompatibility
and biomechanical properties as dental implant
42-44
material has been extensively investigated.
Different materials have been introduced for the
construction of dental implants. Table 6 illustrates
34,35
the mechanical properties of these materials.
The pure titanium (Ti) and Ti alloy offer favorable
physical properties. Ti offers many of the
desirable attributes for an implant material such
as excellent inertness, resistant to corrosion by
body fluids, and apparent compatibility with living
tissue.1,36,37
Hydroxyapatite (HA) ceramic has been
investigated extensively and used for dental
implant applications for the past 30 years. HA
properties depend on its porosity. Clinical
studies demonstrated HA ceramics still remain
the most biocompatible bone implant material
known and possess the added benefit of strong
bonding to living bone through natural-appearing
bonding mechanisms,45 but they can fracture
during surgery and after loading.46 Therefore, HA
ceramics are not ideal materials for permanent
implant devices. However, bioactive ceramic
coatings on metal implants has helped to
retain ceramics as a key component in dental
47,48
implantology.
Ceramic material like aluminum oxide (Al2O3)
38
has been utilized as a dental implant material.
This material has osseointegrated favorably,
but unfortunately the biomechanical properties
(modulus 300GPa) of the implants were not
sufficient for a long-term load. As a result, this
material has been withdrawn from the market as
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The Journal of Contemporary Dental Practice, Volume 9, No. 2, February 1, 2008
DC% level. However, further studies are needed
to determine the surface characteristics required
for optimizing the bonding of FRC to bone by using
such strategies as bioactive ceramic coatings on
FRC implants.
FRC are durable materials with good flexural
properties because it has a lower elastic modulus
than metals.22,49 In fact, the mechanical properties
and modulus of elasticity of unidirectional FRC
(20-40 GPa) are close to natural bone. The FRC
structure makes it possible to embed bioactive
ceramic particles (such as BAG) within the resin
matrix or apply them directly to the surface of
FRC implants. Such an implant structure may be
suitable for the clinical applications involving poor
bone quality.
It is well known higher DC% results in better
mechanical properties for the polymer and
composite material than lower DC%.57 A
higher DC% can be obtained by lengthening
the polymerization time and increasing the
polymerization temperature. Improvement in
mechanical properties by the increased DC%
are related to the increased cross-linking
density which increases the number of covalent
bonds between polymer backbones and lowers
the amount of residual monomers that could
58-60
plasticize the polymer matrix.
The present
study also confirmed this by showing substantial
improvement in bending and torsional properties
of the specimens with a high DC%. A three times
higher load to failure ratio of implant specimens
was observed in group D compared to group A.
The quantity of leachable residual monomers
decreases with an increasing DC%. Thus,
in addition to improved strength, a high DC%
obviously improves the biocompatibility which is
required for implant materials.
The most typical application of FRC material
in dentistry is on endodontically treated teeth
restored with fiber posts. The published
information regarding the clinical survival rate
failure rate was reported to be 12.8% within a 24
month observation period after fiber reinforced
50
posts were placed. While another study showed
the cumulative failure rate for metal posts was
11.2% within ten years.51 However, the typical
difference between metal and FRC posts in
cases of failure is metal post causes root fracture
whereas failure of FRC post occurs at the crown
level leaving the root intact and repairable. This
phenomenon can be explained by the mismatch
between the high stiffness metal and the dentin,
thus, supporting the concept of an FRC implant.
The fatigue test showed statistically significant
differences among the different posts due to
variations in fiber post manufacturing process.52
Commercially available prefabricated FRC
posts showed lower flexural properties than an
individually polymerized FRC material.53
In groups A-D the failure occurred by producing
cracks and delamination of resin prepregs
(interfacial failure), followed by a growing crack at
the interface between the prepregs. This problem
could be overcome by improving the interfacial
adhesion of the prepregs used in specimen
fabrication or by alternatively using a different
fabrication process as was used in groups E and F.
This study was conducted to determine some
mechanical properties of a metal-free FRC
material which could be further evaluated as a
potential dental implant material. The test set-up
was designed to simulate simple clinical loading
conditions of an implant.
Specimens in group E and F were prepared to
avoid delamination by changing the fabrication
process. In group E the specimens were coated
with fiber weave to encapsulate the unidirectional
fibers together (Figure 6), whereas in group F the
specimens were made of manually impregnated
fibers as one large sized roving. The four times
higher load to failure ratio obtained in group E
compared to Group A suggests the fiber weave
effectively protected the unidirectional fibers
against delamination.
The introduction of a new biomaterial for surgical
use requires a number of preclinical studies.
The materials used in this study were made
of pBis-GMA-pTEGDMA copolymer, which is
already widely used and generally accepted as
54-56
bone cement.
From this perspective, the
currently tested resin materials can be estimated
to be acceptable for the use in dental implants,
especially if they have been polymerized to a high
This in vitro
o study showed the mean static load
to failure ratio of specimens in all groups was
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The Journal of Contemporary Dental Practice, Volume 9, No. 2, February 1, 2008
torque data was not available for the specimens
in groups A-D. Encapsulation of unidirectional
FRC with a fiber weave (Group E) improved
the resistance to torsion loading. Since this
specimen design showed the best mechanical
properties in the present study it seems to hold
the greatest potential for in vivo
o studies.
The basic challenge in the manufacture of a FRC
implant has been reinforcement of the thread on
the screw shape implant design which has been
overcome by reinforcing the threads with external
bidirectional fiber coating.61
Figure 6. Schematic diagram of test specimen
fabricated for group E: continuous unidirectional FRC
core (A) was encapsulated with a bidirectional fiber
weave (B) (0 +/-90° fiber angle).
In a clinical situation FRC implant fractures offer
a challenging problem because of functional,
restorative, and surgical implications. This
complication can be solved by restoring the FRC
implant with composite superstructures.
above the previously reported maximum incisal
biting force. Whereas the mean load of failure
for groups D, E, and F were beyond the reported
maximum posterior bite force.13 However, flexural
fatigue studies are needed to simulate the
dynamic loading conditions by the masticatory
system.
Conclusions
Within the limitations of this study it can be
concluded:
1. Failure forces of FRC specimens with an
average diameter of dental implant exceeded
reported maximum static human bite forces.
2. Encapsulation of continuous unidirectional
glass fibers with a fiber weave, improves the
resistance FRC specimens to torsional and
bending forces.
3. Prolonging of polymerization time and
increasing the polymerization temperature
increases the DC% of the polymer matrix of
FRC.
Torsional loading produced maximum shear
stress and twisted the structure along the neutral
axis. Shear stress increases as a function of
distance from the neutral axis. With a torsional
load, the specimens in groups A, B, C, and
D exhibited twisting at the very early moment
of loading and delamination occurred before
a torsional load was registered. As a result,
References
1. Brånemark PI, Hansson BO, Adell R, Breine U, Lindstrom J, Hallen O. Osseointegrated implants in
the treatment of the edentulous jaw: experience from a ten year period. Scand J Plast Reconstr Surg
Suppl. 1977; 16:130-6.
2. Jemt T, Linden B, Lekholm U. Failures and complications in 127 consecutively placed fixed partial
prostheses supported by Branemark implants. Int J Oral Maxillofacial Implants. 1992; 7:40-4.
3. Vidyasagar L, Apse P. Biological response to dental implant loading/overloading. Implant
overloading: Empiricism or science. Stomatologija. 2003; 5:83-9.
4. Rangert B, Krogh PH, Langer B, Van Roekel N. Bending overload and implant fracture. Int J
Oralmaxillofac Implants. 1995; 10:326-34.
5. Hiermer T, Schmitt-Thomas KH, Yang Z. Mechanical properties and failure behavior of cylindrical
CFRP-implant-rods under torsion. Composites. 1998; 29:1453-61.
6. Richter EJ. In vivo
o horizontal bending moments on implants. Int J Oral Maxillofac Implants. 1998;
13:232-44.
7. Mao JJ, Major PW, Osborn JW. Coupling electrical and mechanical outputs of human jaw muscles
undertaking multidirectional bite-force tasks. Arch Oral Biol. 1996; 41:1141-7.
8. Morgan MJ, James DF. Force and moment distributions among osseointegrated dental implant.
J Biomechanics. 1995; 28:1103-9.
9
The Journal of Contemporary Dental Practice, Volume 9, No. 2, February 1, 2008
9. Weinberg LA. The biomechanics of force distribution in implant-supported prostheses. Int J Oral
Maxillofac Implants. 1993; 8:19-31.
10. Simon JA, Ricci JL, Di Cesare PE. Bioresorbable fracture fixation in orthopedics: a comprehensive
review. Part I. Basic science and preclinical studies. Am J Orthop. 1997; 26:665-71.
11. Laurell L. Occlusal forces and chewing ability in dentitions with cross-arch bridges. Swed Dent J
Suppl. 1985; 26:160-7.
12. Paphangkorakit J, Osbom JW. The effect of pressure on maximum incisal bite force in man. Arch
Oral Biol. 1997; 42:11–7.
13. Van Eijden TM. Three-dimensional analyses of human bite-force magnitude and moment. Arch Oral
Biol. 1991; 36:535-9.
14. Smith DC. The acrylic denture: mechanical evaluation of midline fracture. Br Dent J. 1961;
110:257-67.
15. Freilich MA, Duncan JP, Alarcon EK, Eckrote KA. The design and fabrication of fiber-reinforced
implant prostheses. J Prosthet Dent. 2002; 88:449-54.
16. Behr M, Rosentritt M, Lang R, Handel G. Glass fiber-reinforced abutments for dental implants.
A pilot study. Clin Oral Implants Res. 2001; 12:174-8.
17. Duncan JP, Freilich MA, Latvis CJ. Fiber-reinforced composite framework for implant-supported
overdentures. J Prosthet Dent. 2000; 84:200-4.
18. Tanner J, Vallittu PK, Soderling E. Effect of water storage of E-glass fiber-reinforced composite on
adhesion of Streptococcus mutans. Biomaterials. 2001; 22:1613-8.
19. Tanner J, Robinson C, Soderling E, Vallittu PK. Early plaque formation on fiber-reinforced
composites in vivo. Clin Oral Investig. 2005; 9:154-60.
20. Vallittu PK. Some aspects of the tensile strength of unidirectional glass fiber-polymethylmethacrylate
composite used in dentures. J Oral Rehabil. 1998; 25:100-5.
21. Chung K, Lin T, Wang F. Flexural strength of a provisional resin material with fiber addition. J Oral
Rehabil. 1998; 25:214-7.
22. Chong KH, Chai J. Strength and mode of failure of unidirectional and bidirectional glass fiberreinforced composite materials. Int J Prosthodont. 2003; 16:161-6.
23. Goldberg AJ, Burstone CJ. The use of continuous fiber reinforcement in dentistry. Dent Mater. 1992;
8:197-202.
24. Vallittu PK. The effect of void space and polymerization time on transverse strength of acrylic-glass
fiber composite. J Oral Rehabil. 1995; 22:257-61.
25. Loza-Herrero MA, Rueggeberg FA, Caughman WF, Schuster GS, Lefebvre CA, Gardner FM. Effect
of heating delay on conversion and strength of a post-cured resin composite. J Dent Res. 1998;
77:426-31.
26. Vallittu PK. Effect of 180-week water storage on the flexural properties of E-glass and silica fiber
acrylic resin composite. Int J Prosthodont. 2000; 13:334-9.
27. Lassila LV, Nohrstrom T, Vallittu PK. The influence of short-term water storage on the flexural
properties of unidirectional glass fiber reinforced composites. Biomaterials. 2002; 23:2221-9.
28. Tuusa SM, Lassila LV, Matinlinna JP, Peltola MJ, Vallittu PK. Initial adhesion of glass fiber-reinforced
composite to the surface of porcine calvarial bone. J Biomed Mater Res B Appl Biomater. 2005;
75:334-42.
29. Ballo AM, Kokkari A, Meretoja V, Lassila LV, Vallittu PK, Närhi TO. In vitro
o biocompatibility evaluation
of fiber-reinforced composite implant device. J Mater Sci-Mater Med. (Submitted).
30. Aldini NN, Fini M, Giavaresi G, Torricelli P, Martini L, Giardino R, Ravaglioli A, Krajewski A,
Mazzocchi M, Dubini B, Ponzi-Bossi MG, Rustichelli F, Stanic V. Improvement in zirconia
osseointegration by means of a biological glass coating: An in vitro
o and in vivo
o investigation.
J Biomed Mater Res. 2002; 61:282-9.
31. Moritz N, Rossi S, Vedel E, Tirri T, Ylanen H, Aro H, Narhi TO. Implants coated with bioactive glass
by CO2-laser, an in vivo
o study. J Mater Sci Mater Med. 2004; 15:795-802.
32. Pastila P, Lassila LV, Jokinen M, Vuorinen J, Vallittu PK, Mantyla T. Effect of short-term water
storage on the elastic properties of some dental restorative materials-A resonant ultrasound
spectroscopy study. Dent Mater. 2007; 23(7):878-84.
10
The Journal of Contemporary Dental Practice, Volume 9, No. 2, February 1, 2008
33. Lum LB, Osier JF. Load transfer from endosteal implants to supporting bone: an analysis using
statics. Part two: Axial loading. J Oral Implantol. 1992; 18:349-53.
34. O’Brien WJ. Dental materials and their selection. Quintessence Publishing Co, Inc. 2002:340-80.
35. Bouillaguet S, Schutt A, Alander P, Schwaller P, Buerki G, Michler J, Cattani-Lorente M, Vallittu PK,
Krejci I. Hydrothermal and mechanical stresses degrade fiber-matrix interfacial bond strength in
dental fiber-reinforced composites. J Biomed Mater Res B Appl Biomater. 2006; 76:98-105.
36. Adell R, Lekholm U, Rockler B, Branemark PI. A 15-year study of osseointegrated implants in the
treatment of the edentulous jaw. Int J Oral Surg. 1981; 10:387-416.
37. Parr GR, Gardner LK, Toth RW. Titanium: the mystery metal of implant dentistry. Dental materials
aspects. J Prosthet Dent. 1985; 54:410-4.
38. Schulte W. The intra-osseous Al2O3 (Frialit) Tuebingen Implant. Developmental status after eight
years (I-III). Quintessence International 1984; 15:1-39.
39. Drouin JM, Cales B. Yttria-stabilized zirconia for improved hip joint head. In: Andersson ÖH,
Yli-Urpo A, editors. Bioceramics 7. London: Butterworth-Heinemann Publ.1994:387-94.
40. Piconi C, Burger W, Richter HG, Cittadini A, Maccauro G, Covacci V, Bruzzese N, Ricci GA,
Marmo E. Y-TZP ceramics for artificial joint replacements. Biomaterials. 1998; 19:1489–94.
41. Richter HG, Burger W, Osthues F. Zirconia for medical implants-the role of strength properties.
In: Andersson OH, Yli- Urpo A, editors. Bioceramics 7. London: Butterworth-Heinemann Publ.,
1994:401-6.
42. Albrektsson T, Hansson HA, Ivarsson B. Interface analysis of titanium and zirconium bone implants.
Biomaterials. 1985; 6:97–101.
43. Ichikawa Y, Akagawa Y, Nikai H, Tsuru H. Tissue compatibility and stability of a new zirconia ceramic
in vivo. J Prosthet Dent. 1992; 68:322–26.
44. Kohal RJ, Klaus G, Strub JR. Zirconia-implant-supported all-ceramic crowns withstand long-term
load: a pilot investigation. Clin Oral Implants Res. 2006; 17:565-71.
45. Chang YS, Oka M, Nakamura T, Gu HO. Bone remodeling around implanted ceramics. J Biomed
Mater Res. 1996; 30:117-24.
46. Jarcho M. Retrospective analysis of hydroxyapatite development for oral implant applications.
Dent Clin North Am. 1992; 36:19-26.
47. Krauser JT, Boner C, Boner N. Hydroxyapatite coated dental implants. Biological criteria and
prosthetic possibilities. Cah Prothese. 1990; 71:56-75.
48. Ogiso M, Nakabayashi N, Matsumoto T, Yamamura M, Lee RR. Adhesive improvement of the
mechanical properties of a dense HA-cemented Ti dental implant. J Biomed Mater Res. 1996;
30:109-16.
49. Lassila LV, Nohrstrom T, Vallittu PK. The influence of short-term water storage on the flexural
properties of unidirectional glass fiber-reinforced composites. Biomaterials. 2002; 23:2221-9.
50. Naumann M, Blankenstein F, Dietrich T. Survival of glass fibre reinforced composite post
restorations after 2 years-an observational clinical study. J Dent 2005; 33:305-12.
51. Balkenhol M, Wöstmann B, Rein C, Ferger P. Survival time of cast post and cores: a 10-year
retrospective study. J Dent 2007; 35:50-8.
52. Grandini S, Goracci C, Monticelli F, Tay FR, Ferrari M. Fatigue resistance and structural
characteristics of fiber posts: three-point bending test and SEM evaluation. Dent Mater. 2005;
2:75-82.
53. Lassila LV, Tanner J, Le Bell AM, Narva K, Vallittu PK. Flexural properties of fiber reinforced root
canal posts. Dent Mater. 2004; 20:29-36.
54. Evans SL, Hunt CM, Ahuja S. Bone cement or bone substitute augmentation of pedicle screws
improves pullout strength in posterior spinal fixation. J Mater Sci-Mater Med. 2002; 13: 1143-5.
55. Palussiere J, Berge J, Gangi A, Cotten A, Pasco A, Bertagnoli R, Jaksche H, Carpeggiani P,
Deramond H. Clinical results of an open prospective study of a bis-GMA composite in percutaneous
vertebral augmentation. Eur spine J. 2005; 14:982-91.
56. Andreassen GS, Hoiness PR, Skraamm I, Granlund O, Engebretsen L. Use of a synthetic bone
void filler to augment screws in osteopenic ankle fracture fixation. Arch Orthop Trauma Surg.
2004; 124:161-5.
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57. Braga RR, Cesar PF, Gonzaga CC. Mechanical properties of resin cements with different activation
modes. J Oral Rehabil. 2002; 29:257-62.
58. Rueggeberg FA, Caughman WF, Curtis JW Jr. Effect of light intensity and exposure duration on cure
of resin composite. Oper Dent. 1994; 19:26–32.
59. Asmussen E, Peutzfeldt A. Mechanical properties of heat treated composite resin inlay/onlay
technique. Scand J Dent Res. 1990; 98:564-7.
60. Loza-Herrer MA, Rueggeberg FA. Time temperature profiles of post-cure composite oven. Gen Dent.
1998; 46:79-83.
61. Ballo AM, Lassila LV, Vallittu PK, Närhi TO. Load bearing capacity of bone anchored fiber-reinforced
composite device. J Mater Sci-Mater Med. (Accepted for publication).
About the Authors
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Acknowledgement
The first author would like to thank the ITI Foundation for providing a research scholarship in support of
this investigation.
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The Journal of Contemporary Dental Practice, Volume 9, No. 2, February 1, 2008